Freeze–thaw hydrogel fabrication method: basic principles, synthesis parameters, properties, and biomedical applications

Hydrogel is being broadly studied due to their tremendous properties, such as swelling behavior and biocompatibility. Numerous review articles have discussed hydrogel polymer types, hydrogel synthesis methods, hydrogel properties, and hydrogel applications. Hydrogel can be synthesized by physical and chemical cross-linking methods. One type of the physical cross-linking method is freeze-thaw (F–T), which works based on the crystallization process of the precursor solution to form a physical cross-link. To date, there has been no review paper which discusses the F–T technique specifically and comprehensively. Most of the previous review articles that exposed the hydrogel synthesis method usually mentioned the F–T process as a small part of the physical cross-linking method. This review attempts to discuss the F–T hydrogel specifically and comprehensively. In more detail, this review covers the basic principles of hydrogel formation in an F–T way, the parameters that influence hydrogel formation, the properties of the hydrogel, and its application in the biomedical field.


Introduction
Hydrogels are cross-linked polymer with 3D-structure composed of hydrophilic groups and very flexible due to their large water content [1][2][3]. In recent years, studies and research on hydrogels have grown steadily due to their auspicious potential. They have been applied in many medical, engineering, and electronics industries, e.g., drug delivery system, wound healing, tissue engineering, brain tissue, wastewater treatment, batteries, polymer electrolyte, smart supercapacitor, flexible sensors and actuators [4][5][6][7][8][9][10][11][12]. Hydrogels are flexible material, so they are easy to shape, depending on the desired application. Hydrogels can be applied in the form of gel gradients, anisotropic gels, gel patterns, gel wrinkles, nanoparticle gel, and tube structures [13]. The promising potential of hydrogels has attracted many parties to develop fabrication methods to obtain maximum results.
Hydrogels can be classified into physical and chemical methods based on the way in which the cross-links are formed [28]. A more detailed and comprehensive comparison of these fabrication methods is presented in table 1. The physical cross-linking at the molecular scale involves non-covalent interactions, e.g., freeze-thaw (F-T), stereo complex formation, hydrogen bonding, or chain entanglements, while chemical cross-linking is often fabricated by chemical grafting, radiation, and polymerization [29,30]. Hydrogel fabrication techniques that use chemical cross-linking tend to produce stable hydrogels with high mechanical strength, high superabsorbent properties, and are insoluble in all solvents except covalent cross-linking [31]. However, the chemical agents are usually poisonous, so they must be neutralized prior to application [32]. For biomedical applications, chemical aspects must be considered in synthesizing hydrogels. This chemical reaction may cause damage to the biocompatibility of the hydrogel, so the physical cross-linking mechanism could be a promising option.
Physical crosslinking has become a popular method because of the simplicity of the synthesis process. Also, it can produce hydrogels with promising properties without chemical cross-linking agents, thus avoiding potential cytotoxicity from unreacted chemical crosslinkers [33]. Among the physical cross-linking methods, the F-T process is the most broadly used recently. The advantages of the F-T method are that it is simple and it requires no specific equipment for hydrogel synthesis [34]. This method could produce hydrogel with highly biocompatible, high mechanical properties, an excellent swelling ratio, and non-toxic [35][36][37][38].
Numerous review papers about hydrogels have been published with specific issues. Several review papers about hydrogels have been reviewed thoroughly e.g., the raw materials for making hydrogels (synthetic and natural polymers), the properties of the hydrogels, the potential applications of the hydrogels, and the synthesis methods of the hydrogels [39][40][41][42][43]. However, a review article that comprehensively describes the hydrogel formation scheme using the F-T method has yet to be found. Therefore, this review aims to explain the formation mechanism of the hydrogel using the F-T method, the effect of various process parameters on the properties of the developed hydrogel, and its application in several biomedical fields. Moreover, the highlights of the potential for future development of F-T hydrogel are also provided.

Freeze-thaw method
The F-T method is basically a physical crosslinking method without involving chemical covalent bonding agents. The mechanism of physical hydrogel formation can be done through electrostatic schemes, hydrogen bonds, or hydrophobic forces between polymer chains. While this scheme requires a polymer network that meets the following conditions: (1) strong interchain interactions to form stable collocations in the molecular network; and (2) polymer networks must encourage access to and residence of water in the hydrogel [52].
Peppas first introduced the F-T method in 1975, which studied the super-molecular structure of PVA-based hydrogels [53]. By that time, the F-T process was becoming increasingly popular. Willcox (1999) utilized the F-T method to investigate the microstructure of PVA hydrogels [54]. The crystallinity properties of the hydrogels produced by the F-T method have also been described in full [55]. Using the F-T method, hydrogels can be made from natural polymers such as polysaccharides instead of synthetic ones [56]. The high interest in research on F-T hydrogels is evidenced by the increasing number of documents that have been published in the last few decades spread across several subject areas (figure 1). Recent developments in the F-T method are to produce artificial glenoid labrum-based hydrogels, extracellular matrix, wound healing materials, delivery of herbal medicines, flexible supercapacitors, flexible electrodes, heavy metal ion removers, and absorbent materials [34,[57][58][59][60][61][62][63].
2.1. Basic principle of the freeze-thaw method As one of the physical cross-linking methods to fabricate hydrogels, F-T has two stages of hydrogel formation, i.e., the freezing stage of the precursor solution under 0°C, and the thawing stage at room temperature. The main idea is to control the ice crystallization process (freezing) and the formation of an ordered structure (thawing) so that the hydrogels possess optimal properties [64]. Hydrogels are made of polymers which are selected based on their superior properties and applications, e.g., polyvinyl alcohol with biocompatible properties is very suitable for biomedical applications, polyaniline as a conductive polymer is a potential candidate for supercapacitor applications, polyacrylamide can be applied to water treatment because of its ability to thicken suspended solids, and cellulose-based hydrogel for agricultural applications [25,[65][66][67]. Hydorgel had pH sensitivity and non-Fickian drug release. However, the toxic, unreacted monomer residues require a cleaning time of almost 2 weeks.
Heated in water bath at 55°C, increased to 70°C for 24 h and then cooled at RT.
The resulting hydrogel was highly porous and pH sensitive. It also performed the Korsmeyer-Peppas drug release, followed by a non-fickian diffusion mechanism.
Drug delivery. [45] Ionizing radiation -Materials: AgNPs/PVP/PVA. -Crosslinking agent: -Iradiated using 60 Co at 25 kGy with dose level of 1.58 kGy/h. Without a crosslinking agent, it is effective against gram-positive and gram-negative bacteria and their resistance. However, its mechanical properties still need to be investigated further.
Wound dressing [46] One-pot synthesis -Materials: rGO/gelatin. -Crosslinking agent: -Heated in an oil bath at 95°C for 24 h. It is a green and facile method without using chemical crosslinkers or organic solvents, but at high concentrations, graphene can be toxic.
Mixing PDLA and PLLA solutions to obtain a stereocomplex at room temperature. The gelatin solution was added at 0°C and then frozen at −20°C. Lastly, the hydrogel was immersed in the genipin solution for 24 h.
The produced hydrogel featured ideal interconnectivity, abundant porosity, and proper size, which are useful for nutrient supply, cell proliferation, and calcium deposition. However, monitoring degradation and controlling polymer constitution so that the rate of degradation is perfectly consistent with new bone formation is still challenging.
GelMA hydrogels were formed by adding APS and TEMED to the GelMA solution at 37°C for several minutes. After gelation, the hydrogel was immersed in the TA solution for a predetermined time, then the free TA was rinsed with deionized water.
Hydrogel exhibited superior mechanical properties, excellent biocompatibility, and the best recovery results in skin and gastric wounds. However, the strong hydrogen bonding interaction between GelMA and TA can suppress the hydrogel structure thereby limiting its water loading capacity.
The solution was ultrasonicated in an ice water bath for 15 min Then the solution was heated to 70°C with an oil bath for 2-6 h under N 2 atmosphere, and the polymerization process was carried out using APS.
The synthesized hydrogel had a unique swelling behavior, excellent flexibility, and mechanical strength. However, its biocompatibility and biodegradability still need to be investigated further.
The hydrogel was produced without a crosslinking agent, had excellent hydrophilicity, viability up to 295%, was able to reduce bacteria up to 99.94%, and had mechanical properties like skin tissue.

Polymers cross-linking mechanism on hydrogel
Polymers are materials consisting of long chains of molecules with repeating units [68]. The intrinsic structure of polymers, such as bonds, molecular chain arrangement, and molecular weight, influences their properties [69].
Bonds determine the stability of atoms and molecules' bond . Regular and irregular structures affect the strength and toughness of the polymer. Molecular weight influences the viscoelasticity and other physical properties of the polymer. Atoms in individual polymer molecules are connected by strong covalent bonds, whereas intermolecular forces link polymer molecules. During the hydrogel formation process using the F-T method, physical cross-linking occurs due to the crystallization of the polymeric solution.
The cross-linking mechanism is illustrated in figure 2(a). The interaction of the hydroxyl groups in the polymer chain with water molecules causes hydrogen bonds to form. The freezing process of the polymer solution induces crystal growth which acts as a cross-linking point between polymer chains. The thawing treatment relaxes the polymer chains so that they can move freely again. The F-T cycle is repeated continuously causing more and more cross-link points to be formed. The point of crosslinking of the polymer chains eventually forms a three-dimensional matrix which makes the solvent trapped in the matrix, resulting in a gel phase [70].

Hydrogel formation parameters
The produced hydrogel properties are influenced by the synthesis method and the original properties of the precursors. According to Hassan & Peppas [71], the F-T hydrogel properties depend on the molecular weight of the polymer, the solution concentration, F-T temperature, time, and the number of cycles. A summary of various parameters for the synthesis of F-T hydrogels is given in table 2. Although most of the polymers for the F-T method rely on PVA, there are several other polymeric materials such as chitosan, alginate, gelatin, several types of gums, к-carrageenan, and glucomannan that can be made into hydrogels using the same method. Some of these polymers share the same characteristics, which are hydrophilic polymers with many hydrophilic functional groups (-OH, -COOH, -NH 2 ) in the chain structure. The large amounts of hydrophilic groups and the freezing process in the F-T method allows the polymer chains to be bonded non-covalently through the formation of hydrogen bonds (intramolecular, intermolecular I, and intermolecular II) [72]. Furthermore, with this physical bond, the polymer chains can form a 3-dimensional matrix capable of trapping solvents between cross-linking points to produce a gel phase.

The molecular weight of the polymer
The sum of each atomic weight of the atoms making up the molecule is known as the molecular weight (M W ) of the polymer, which indicates the average length of the bulk resin polymer chain [73]. The M W values are expressed in terms of distribution ranges due to their different molecular weights, even the polymers in a particular class [74].
The M W of a polymer is closely related to its viscosity and permeability. Compared to low molecular weight polymers, high molecular weight polymers have a higher permeability reduction factor and greater viscosity [75]. The M W of the polymer has been shown to play an essential role in forming the gel structure [76]. Ma et al have reported the effect of M W variations on the PVA hydrogel properties produced by controlling the entanglement density [77]. The increase in molecular weight led to an increase in the stable hydrogel crosslinked network, gel concentration, and enhanced tensile strength. However, increasing the molecular weight of PVA could decrease the swelling ability of the hydrogel.
As shown in figure 2(b), the variation in molecular weight affects the microstructure of PVA hydrogel [77]. Low molecular weight PVA has a shorter chain length, allowing them to move freely in the solvent. The obtained hydrogel has fewer microcrystalline, cross-linking points, and hydrogen bonds because it is only affected by the interchain. The chain length between the cross-links of low molecular weight PVA is also greater than that of high molecular weight PVA. In high molecular weight PVA, the formation of hydrogen bonds that occur are intrachain and interchain. Thus, the number of cross-linking points, microcrystalline, and hydrogen bonding is more significant [77]. Besides, Xia et al asserted that the PVA molecular weight must be well-determined [78]. PVA with a molecular weight of too low cannot form a cross-linking, while a molecular weight of too high will produce bubbles in the solution that are difficult to remove.

Solution concentration
The F-T method begins with the preparation of a polymer solution. The concentration of a polymer solution is expressed as the weight ratio of the polymer to the solvent. Decreasing the polymer concentration in the solution will decrease the viscosity of the solution and reduce the cross-linking density of the hydrogel, which results in a softer gel [79,80]. Previous studies have shown the effect of variations in solution concentration of the polymer on hydrogel properties. Syifa et al found that with decreasing PVA concentration, the resulting hydrogel became softer but the swelling value increased [80]. Research by Waresindo and Edikresnha showed the same conclusion when loading PVA hydrogel with guava leaf extract and red betel extract [79,81]. Hou and colleagues reported the effect of different concentrations of dimethyl sulfoxide (DMSO) solution on the transparency, mechanical properties, and crystallinity of PVA hydrogels prepared by the F-T method [82]. They found that the increase in DMSO concentration or decrease in PVA resulted in better transparency of the PVA hydrogel. The hydrogen bond between the DMSO molecule and the −OH group of the PVA chain inhibits crystal growth in the twodimensional direction, resulting in a smaller crystal volume. The small crystal volume eases the light penetration and enhances the transparency value. Moreover, Joshi's group examined the effect of PVA solution concentration on its rheological properties during the F-T gelation process [83]. Hence, the PVA hydrogel has better gel strength because increasing the concentration of PVA is aligned with the density of the cross-linked network.

Temperature, duration, and number of F-T cycles
The F-T method relies on a crystallization process to form cross-links in the hydrogel. This crystallization occurs during the freezing process of the polymer solution. The frozen temperature parameters, freezing time, and the repetition of the F-T cycle are the factors affecting the properties of the produced hydrogel.
Kim and colleagues found that various temperatures of PVA solutions during the freezing process could induce the hydrogel to form a stiffness gradient [70]. Low temperatures cause more cross-linking points of the polymer chains to form, so the hydrogel becomes stiffer. Furthermore, Figueroa − Pizano et al reported the swelling ability of chitosan/PVA-based hydrogels at different freezing temperatures (−4, −20, and −80°C) [84]. Based on their study, lower freezing temperatures reduced the pore size of the hydrogel. Thus, the swelling ability decreased due to the increased cross-linking point between PVA chains. Freezing time during the F-T steps can affect the formation of freezable water coagulation.
Nakano and Nakaoki investigated the coagulation size of freezable water in PVA hydrogels during the F-T cycle variations [85]. When the F-T cycle period is short (figure 2(c)), the coagulation of freezable water, the crystallites, and the PVA domain of the amorphous material swells that contain non-freezable water. The cause is the interaction between polymer chains that yields very few cross-linking points. The higher number of the F-T cycle period triggers more cross-linking points in the polymer. Thus, the water molecules in the swollen PVA region and non-freezable water turn into freezable water, enlarging the water domain that can be frozen.
The number of F-T cycles promotes the gelation of PVA solution and increases the mechanical properties of PVA hydrogel. Holloway's group reported the phase separation and crystallization of PVA hydrogels during the F-T process [86]. Samples were frozen for 21 h at 20°C and thawed for three hours at room temperature for up to ten cycles. The compression modulus value increases as the number of F-T cycles increases from cycle one to cycle six. After six cycles, the compression modulus value tends to be stable. Li et al have come to a similar conclusion when observing the effect of the number of F-T cycles on the mechanical and conductive properties of PVA/PANI hydrogel for supercapacitor applications [87].
They found that the F-T process can produce a microstructure such as spongy bone in hydrogel samples that has increased mechanical and electrochemical properties. Their variation of F-T cycles resulted in significant differences in tensile strength values.   During the first cycle, some parts of the chain crystallize, forming major crystal junctions of 3 to 8 nm in size. An irregular mesh separates these crystals by an average distance of 30 nm. [54] PVA/CNC -PVA 10 wt% CNC 0, 5, 10, 20 wt% Freeze at −20°C for 18 h, thaw at RT for 4 h. Repeated for 5 cycles.
The percentage of crystallinity of the hydrogel samples was seen to decrease at all loadings. There is an increase in water absorption when the CNC concentration is increased. The high CNC content means the hydrogel has better structural integrity. Freeze at −20°C for 10 h, thaw at RT for 1 h. Repeated for 0, 1, 3, 5, 7, and 9 cycles.
The increasing number of F-T cycles resulted in a stiffer structure, more stable thermal properties, and a higher degree of crystallinity. In contrast, after 3 F-T cycles, a lamellar structure was formed so that the swelling ratio of the hydrogel equilibrium decreased. [89] Starch/PVA/LRD -Variated of LDR content from 0% to 12.5% of a polymer basis Freeze at −20°C for 10 h, thaw at 25°C for 2 h. Repeated for 3 cycles.
Fast freezing using liquid nitrogen for 20 min, thaw at 4°C for 4 h. Repeated for 2 cycles.
The mechanical properties, crystallinity, and degree of swelling were increased in hydrogels by the orientation method. This increase is due to the alignment of the polymer chains that occurs during the F-T process. The increasing number of cycles causes more polymer alignment to occur.
The SS/PVA hydrogel showed excellent swelling ability and hydrophilicity for its porous structure. PVA blending effectively improved sericin's thermal stability and enhanced sericin's mechanical properties but did not affect sericin and PVA's crystallinity.
[91]  Increasing the CNC concentration enhances the capacity of water absorption and decreases the modulus of compression and optical transparency. [93] PAA/PVA PVA: 85,000-124,000 g mol −1 10 wt% PVA with variations in PAA concentration.
Freeze at −20°C for 12 h, thaw at RT for 6 h. Repeated for 3 cycles.
The variation in acrylic acid concentration affects the swelling and thermal properties of the PAA/ PVA hydrogels. The increase in acrylic acid concentration resulted in increased stability, swelling, and thermal transparency but decreased optical transparency. [94] PVA 146,000-186,000 g mol −1 9% w/v Freeze at −23°C for 16 h, thaw at 25°C for 8 h. Repeated for 3 cycles.
The variation in the number of F-T cycles affects the physico-chemical and viscoelastic properties of the hydrogels. Increasing the number of F-T cycles causes an increase in the degree of crystallinity. Hydrogel crystallinity is closely related to its ability to absorb water. [95] PVA/GLE -PVA: 89,000-98,000 g mol −1 10 wt% PVA with variations in GLE composition.
Freeze at −25°C for 20 h, thaw at RT for 4 h. Repeated for 6 cycles.
An increase in GLE concentration causes an enhance in pore size, degree of swelling, and weight loss. The CS content, the number of F-T cycles, and the freezing temperature significantly affected the degree of swelling. Pore size increases with CS content, but lower temperatures or longer freezing times result in higher porosity and smaller pores. [84] Alginate 80,000-120,000 Da 0.5 and 1 wt% Freeze at -25°C for 24 h and thaw at 4°C for 24 h. Repeated for 1, 2, 3, 4, 5 and 6 cycles.
The F-T treatment increased the hydrogel storage modulus by almost 100 times. The properties of the resulting alginate gel, such as dynamic modulus and gel syneresis, were influenced by the pH value, number of F-T cycles, alginate concentration, and ionic strength. Because of its soft structure and melting behavior during storage, it has the potential for new applications in biomedical fields.
[96] The combination of fast freezing and slow thawing treatment conditions resulted in a uniform polymer network structure after several F-T cycles. The gelatin hydrogel showed an optimized fine inner structure (pore size = 23.72 ± 20.61 μm 2 ), stable latent heat of fusion (248.8 J g −1 ), constant total moisture content (about 90%), high cooling efficiency, and relatively stable mechanical strength (46 kPa).
Compared to curdlan gels formed with guar, кcarrageenan, or locust bean gum, the curdlan/ xanthan composite hydrogel proved to be the most resilient, exhibiting zero syneresis with superior consistency in viscosity, heat stability, storage and loss modulus, gel strength, and adhesion. [98] Curdlan gum1. 9  The presence of zein strengthens the effect of F-T treatment on increasing G′ gel. The F-T treatment resulted in a coarse network structure, increased the G′ of the composite gel, partially destroyed the original crystalline domain of the composite gel, and increased the initial degradation temperature.

Properties of hydrogel
The crucial properties of hydrogels, e.g., morphology and porosity, swelling ability, mechanical properties, thermal properties, and biocompatibility, have become the focus of research in recent decades. These properties can be adjusted based on the requirements of the desired application, selection of appropriate materials, and synthesis methods.

Morphology and porosity
The bulk hydrogel structure is a solid polymer that has absorbed water and has nano-sized pores in the network [101]. Because pore size modulation and distribution are critical for controlling hydrogel properties, they have been widely used in tissue engineering, wound dressings, and drug delivery systems [102][103][104][105]. The hydrogels porosity is formed by the formation of cross-links between polymer chains [106]. The pore diameter (d) can be classified into nanopores, micropores, and macropores, which are < d nm , respectively [107]. Figure 3(a) shows the non-porous hydrogels matrix that can be in the form of a tight cross-link, and there are only small pores with a size of tens of nanometers, for example, the alginate network [108,109]. The hydrogel matrix can also have macroscopic pores with sizes in the range of 10-500 μm [110]. It has been mentioned in the previous section that the hydrogel pore size can be influenced by the parameters of the hydrogel formation process.
Process parameters that increase the number of cross-linking points will make the hydrogel stronger, but the pore size will decrease. Several methods have been developed to adjust the pore size of the hydrogel, for example, gas foaming, reverse casting, particle leaching, and freeze-drying [111]. Among the several methods mentioned above, freeze-drying is the most widely used. The main principle of the freeze-dry way is a fast-cooling process, which creates thermodynamic instability, resulting in phase separation. The solvent content in the hydrogel is sublimated in a vacuum, leaving a cavity in the previously occupied area [112].
Hydrogel porosity can be measured using the liquid displacement method assisted by a solvent that does not react with hydrogel, water, or ethanol/methanol [113]. The solvent used must be able to reach the entire void left to properly map the pores of the hydrogel. As a result, the size of the solvent molecules has a significant impact on the accuracy of the hydrogel porosity. Tao et al reported the equation to calculate hydrogel porosity as follows [91]: where V 1 in equation (1) refers to the initial water volume, V 2 recorded as the total volume of water when the hydrogel is put into the beaker, and V 3 refers to water volume after the hydrogel is removed. Freeze-thawed hydrogels with a macropores structure previously have been investigated by Waresindo et al [79]. Other research conducted by Li et al demonstrated a high porosity hydrogel using agarose (AG) as a poreinducing agent [114]. As given in figure 3(b), the AG transformed the structure of the PVA matrix from initially dense to becoming porous. The addition of AG concentration enhanced the hydrogels' pore diameter (20-200 μm).
The percentage of porosity calculated using equation (1) showed that the levels of porosity varied from 56% (pure PVA), 69% (AG 2 wt%), 78% (AG 4 wt%), and 81% (AG 6 wt%). In this case, AG not only initiated pore formation by initiating nucleation and growth of ice particles, but it also improved mechanical properties by increasing PVA crystallinity and hydrogen bonding between molecules. The pores in the hydrogel promote cell growth, vascularity, and nutrient diffusion. The increase in pore size aids the efficiency of angiogenesis in the hydrogel, provides faster swelling kinetics, and improves absorbency compared to non-porous hydrogels [115].

Swelling ability
The most exciting property to study in hydrogels is their swelling ability. Swelling is a continuous transition from an unbreakable or partially rubber-like state of glass to a relaxed rubber-like state [116]. When the hydrogel meets the solution, it begins to absorb the liquid, which allows the hydrogel to expand. Hydrogels have the ability to absorb water and expand up to 1000 times their dry weight [101]. In addition to the parameters mentioned earlier, the swelling degree and the change in hydrogel dimensions can also be influenced by the hydrophilic/ hydrophobic balance of the hydrogel, the degree of cross-linking, the level of ionization, and its interaction with the counterion [117].
The swelling mechanism of the hydrogel is shown in figure 3(c) [43]. The swelling of the hydrogel works based on the principle of osmotic pressure (the difference in the concentration of mobile ions present in the dry hydrogel structure and then moves to the surrounding solution). The swelling mechanism is also influenced by changes in the acidity of the solvent, which causes changes in the level of ionization of functional groups so that the hydrogel volume becomes larger [118].
The swelling ratio of hydrogels can be determined using the gravimetric method. García-Astrain et al have reported the procedure for measuring the swelling ratio value of the alginate-based hydrogel by immersing the freeze-dried hydrogels into three different solutions [120]. The equilibrium swelling was assumed to be reached when the weight of the hydrogel was no longer increased. By using equation (2), the swelling ratio value can be determined as follows: where W d is the weight of the freeze-dried hydrogels, while W s corresponded to the weight of the swollen hydrogels [120].
The gravimetric method can also be used to calculate the degradation of the hydrogel, which is expressed as weight loss. After measuring the hydrogel weight to determine the swelling value, the sample was re-dried until the weight was stable. Then, by using equation (3), we can determine the percentage of weight loss [121].
Initial weight Weight after degradation Initial weight 100% 3 Thangprasert et al in their research, have successfully fabricated gelatin/PVA hydrogels using the F-T method for tissue engineering applications [119]. Further characterizations of the swelling ratios and weight loss of the hydrogel were investigated, and the results are depicted in figure 3(d). The swelling ability of the hydrogels was observed from one to 180 min. In general, all samples showed a significant swelling rate ranging from 1 to 30 min, then reached swelling equilibrium (no longer increasing in weight). Various compositions of gelatin and PVA yielded different impacts on the swelling ability of the hydrogels. Hydrogels with a gelatin/PVA composition of 100:0 and 70:30 had a higher swelling ratio than the other samples. Sample 0:100, which is pure PVA, showed the lowest swelling value because it has a dense hydrogel matrix and a higher cross-linking degree. The ability of PVA to form -OH bonds between the chains will decrease due to the addition of gelatin concentration, which increases the pore size [92].
Fan and colleagues also observed the swelling ratio of PVA/chitosan/gelatin hydrogel with chitosan/gelatin concentration ratios ranging from 1: 3, 1: 2, 1: 1, 2: 1, to 3: 1 [122]. They found that the capacity of the hydrogel to absorb water was 20-40 times its weight. The swelling rate of the hydrogel gradually increased after 6 to 12 h of immersion and began to slow down after 24 h. The equilibrium swelling state was achieved after 48 h of immersion. Increasing the chitosan concentration caused the hydrogel swelling ratio to increase because chitosan can cause the hydrogel structure to become looser, and the macromolecular chains in the system can be more easily extended. The developed hydrogel then has large pores, and the swelling ratio will be greater too.
Hydrogel degradation is expressed in weight loss percentage, as given in figure 3(d). The cross-linking in PVA is very important in maintaining the physical shape of the hydrogel. This is evidenced by the measurement of weight loss for the other three samples (50:50, 30:70, and 0:100), which retained their shape during the degradation test [119]. Sittiwong's group, who have observed the effect of crosslinking ratio variations on PVA hydrogels, showed a similar conclusion [123]. Increasing the cross-linking ratio caused the hydrogel to be denser and decreased the porosity. The results of the observation of the swelling and weight loss properties also showed a decrease in value with an increase in the concentration of the cross-linking.

Mechanical properties
The mechanical properties of hydrogels play a key role in biological performance. Among the most critical parameters in hydrogel design, the structure and conformation of hydrogels greatly influence their mechanical properties as well as cell adhesion, proliferation, and differentiation [124]. The mechanical properties of the hydrogel include compressive strength, ultimate tensile strength, and modulus of elasticity [64,125].
The best hydrogel mechanical properties are made of synthetic polymers (chemical bonds are stronger than natural polymers) and produced by chemical cross-linking [33,126]. However, with physical treatment (e.g., F-T or anneal swelling) and raw material from natural and synthetic polymers mixture, it has also been reported to yield a hydrogel with good mechanical properties [26].
The tensile stress-strain curves in figure 4(a) indicate a significant increase in the value of the elongation at break and maximum stress when poly (ionic liquid) concentration is increased. Increasing mechanical properties are associated with the increasing number of hydrogen bonds formed between poly (ionic liquid) and PVA [128]. The tensile force can be absorbed by hydrogen bonds, resulting in better mechanical properties [129]. It is proven that the LAPH3 sample with the highest concentration of poly (ionic liquid) had a maximum stress of 0.87 MPa and an elongation at break of up to 265.92%.
In line with the tensile test results, the compressive strength test in figure 4(b) indicates that as poly (ionic liquid) increases, the compressive stress of LAPH also increases. The compressive stress of the LAPH3 sample was 1.61 MPa, while the compressive strain was 7.83%. Furthermore, LAPH3 can also withstand the car's pressure and maintain its original shape, as shown in figure 4(e). The LAPH3 strip is also so elastic that it can stretch up to twice its length (figure 4(f)) and is very soft and easy to tie and stretch ( figure 4(g)). When attached to the hand, as shown in figure 4(h), LAPH3 can remain adhered to without falling due to its stretchability and specific viscosity values.
The mechanical properties of the FT hydrogel can also be tuned by designing the dual network of hydrogel for potential biomedical applications [130]. This double network (DN) method was previously introduced by Gong et al and produced a very high mechanical strength of the hydrogel [131]. DN hydrogels are usually made from a network of stiff, short-chain polymers and then interconnected with a soft and flexible polymer network [132].
Compared with conventional hydrogels, DN hydrogel combinations synergize with each other to produce a hydrogel that is more robust and capable of efficiently dissipating energy throughout the network when subjected to external forces [133]. These advantages make DN hydrogels very potent for tissue engineering applications and drug delivery. However, to be of concern, DN hydrogels that are intended for tissue engineering must have biocompatibility properties that support cell growth, migration, differentiation, and proliferation.

Thermal properties
The characterization of hydrogel thermal properties aims to investigate the state of the polymer and evaluate the interactions between polymer molecules in the hydrogels [89]. The measurement of the thermal properties can be done using a differential scanning calorimetry (DSC). The resulting DSC curves allow us to determine the glass transition temperature (T g ), the melting temperature (T m ), the temperature of crystallization (T c ), and calculate the degree of crystallinity (X c ) of the hydrogel sample [65,134]. The degree of crystallinity (X c ) can be calculated using equation (4) below, where DH f is the melting enthalpy, and DH f 0 refers to the melting enthalpy of a fully crystalline polymer (for PVA 138,60 J g −1 ) [135]. Chee and co-workers have published their work, focused on the effect of orientation on the properties of F-T PVA hydrogel [90]. Their hydrogel samples were prepared with different F-T cycles and uniaxial orientation cycles (100% stretching strain per cycle). The analysis of thermal properties was carried out using the DSC2920 modulated DSC from the TA instrument. Samples were weighed around 7-10 mg with a ramp heating mode from 20°C to 280°C and a heating rate of 10°C min −1 under nitrogen gas flow. The DSC curve of PVA hydrogels with different F-T cycles and orientation is shown in figure 5(a).
The T g of PVA was seen at the peak of 40-64°C, which was represented by relaxation a. The solvent cast PVA sample had a T g of 63.84°C, then when the sample was freeze-thawed, the temperature decreased to 39.93°C. According to Nugent and co-workers, the plasticization effect due to the presence of water in the hydrogel sample is the main cause of this decrease in T g value [136]. The relaxation in the PVA crystal domain (βrelaxation) causes the presence of broad peaks observed at temperatures around 140°C due to the evaporation of the remaining water in the hydrogel sample [137].
The melting temperature (T m ) of the PVA crystal domain was observed at a peak of 223°C which was relatively large and sharp [138]. Increasing the FT cycle and orientation cycle causes an increase in the endothermic PVA curve to become sharper and a shift of the peak to a lower temperature. In line with Chee's group, the thermal properties of PVA hydrogels have also been reported by Butylina et al [134]. They made PVA hydrogels/cellulose nanocrystals (CNC) using the F-T technique with varying numbers of cycles and cellulose concentrations. Thermal properties were analyzed by weighing the sample mass and then characterized using a Mettler Toledo DSC821e tool. The characterization results are listed in table 3.
Hydrogels samples with cycle variations of three times and five times both showed an increase in their melting temperature when the PVA concentration was increased from 5 to 7.5%, then when the PVA concentration was increased to 10%, the melting temperature stopped increasing and might even decrease. The addition of CNC decreased the melting point of all samples except for 7.5% PVA hydrogel samples made with three F-T cycles. These results have been confirmed in the previous studies by Roohani et al and Abitol et al [88,139]. The decrease in melting point is caused by the strong interaction between the polymer matrix and the cellulose surface, limiting the formation of large polymer crystal domains by the polymer matrix chain. The degree of crystallinity is calculated using equation (4), and the results are shown in figure 5(b).
When the PVA concentration increased from 7.5 to 10%, the degree of crystallinity stopped increasing or began to decrease. This is because inhibition of polymer chain folding and crystal formation due to large polymer chain entanglement increases when the polymer concentration is increased [137]. The addition of CNC increased mutual interactions compared to independent associations, resulting in a decrease in PVA crystallinity. A decrease in hydrogel crystallinity due to differences in the number of F-T cycles has previously been reported by Ricciardi et al [55]. They found that the hydrogel crystallinity increased with the increasing number of F-T cycles. However, the crystallinity was relatively static after five or more cycles.

Biocompatibility
For biomedical applications, hydrogels must have biocompatibility properties. Naahidi et al defined biocompatibility as 'a biomaterial's ability to perform with an appropriate host response in the specific application' [140]. Biocompatibility consists of biosafety and bio-functionality parameters. Biosafety is the local and systemic response of the host to the presence of the applied material (the material must be non-toxic, nonmutated, and non-carcinogenic). On the other hand, bio-functionality is the ability of a material to function as expected in a particular application [43]. The biocompatibility of hydrogels can be evaluated through the hematolysis test, systemic allergy test, pyrogen test, conjunctival stimulation test, toxicity test, and cell growth [141]. It has been mentioned before that the hydrogel is made using F-T without involving toxic cross-linking agents, so it can be said that this hydrogel is very safe to use for biomedical applications [79,81,103].
The cytotoxicity test of hydrogel samples can be carried out by in vitro cell viability using the 3-(4, 5-dimethyl2-thiazolyl) -2, 5-diphenyl-2H-tetrazolium bromide (MTT) assay [142]. Cytotoxicity levels can be classified based on the percentage of cells that survived during the MTT assay, also known as the relative growth rate (RGR), which can be calculated using the following equation (5). [143].

( ) =
Optical density value of sample Optical density value of negatif control RGR 5 The RGR percentage values greater than 100, 75-99, 50-74, and 25-49 can be classified as non-poisonous, slightly poisonous, moderately poisonous, and severely poisonous, respectively. Liu et al have successfully reported a double network polyvinyl alcohol/gelatin/glycerol (PVA/GEL/GL) organohydrogel made using the F-T technique [144]. The samples prepared were hydrogel PVA/GEL/GL, PVA/GL, and GEL/GL by freezing at −25°C for 12 h and thawing at room temperature for two hours.
The procedure for cytotoxicity analysis using the MTT assay was started by culturing the control group, the blank group, and the experimental group of the three types of hydrogel samples, each consisting of 4 samples, in a cell culture incubator at 37°C with 5% CO 2 for 24 h. Then L929 mouse fibroblasts in the logarithmic growth phase and good conditions were installed on 96-well cell culture plates at a cell volume of 5 × 103 cells/well, and the medium content per well was 200 μl. Subsequently, the 96-well plate after 24 h of culture was replaced with an organohydrogel extract solution, which had a concentration of 0.2 g ml −1 , and incubation was continued for 24 h. 20 μl of MTT (5 mg ml −1 in PBS) was added to each well of a 96-well plate in a dark place, then incubated for 4 h at 37°C in a CO 2 incubator. After that, the MTT and culture solution were discarded.
The interaction between cell culture and MTT will produce purple formazan. Therefore, 150 μl dimethyl sulfoxide (DMSO) is used, which can dissolve this formazan. The sample solution was then cultured for 24 h, and an optical density (OD) measurement was carried out, but previously 96-well plates were shocked for 10 min. After the OD value is obtained, the RGR percentage/cell viability can be calculated using equation (5). Cell morphology and the percentage of cell viability are shown in figure 6. Figure 6(a) shows L929 cells incubated for 24 h and 48 h. All samples showed good viability, but there was an increase in the number of L929 cells when incubated for 48 h. In figure 6(b), after 24 h of incubation, L929 cells in all samples showed a cell viability percentage above 90%. However, after 48 h of incubation, only L929 cells in the PVA/GL organohydrogel extract solution showed a decrease in the percentage of cell viability, which was below 90%. These characterization results indicated that the developed organohydrogel has good biocompatibility for application to the human body [144].
To give a more detailed comparison in section 3, we have summarized several studies regarding the characterization of the hydrogel properties produced by the F-T method and their applications, which are presented in table 4.

Biomedical application of F-T hydrogel
Hydrogel has the advantages of increased biocompatibility, tunable biodegradability, precise mechanical strength and a porous structure compared to other types of biomaterials [152]. The F-T method is basically a physical crosslinking mechanism without involving toxic crosslinking agents like dialdehyde, formaldehyde, or tetramethylethylenediamine, etc. These toxic properties need to be avoided in order to be suitable for application in the biomedical field. In addition, the need for biomedical applications that require hydrogels with high mechanical properties can also be provided by the F-T method by optimizing the parameters of the synthesis process. The F-T method has also been proven to produce hydrogels that can load drugs, antioxidants, or antibacterial agents, making them very useful in the biomedical field. In this section, we discuss the application of F-T hydrogel in several important biomedical fields, e.g., wound dressing, drug delivery systems, and tissue engineering.

Wound dressing
Wound dressings are usually made of polymers in the form of gauze, gel, hydrogel, or hydrocolloid. The hydrogel is the most promising form and is ideal for application as a wound dressing because it keeps moisture in the wound area, helps remove wound exudate, prevents infection, and facilitates tissue regeneration [153]. A wound is damage that occurs on the inside or outside of the body's tissues, which can result from a collision, cut, or other damage. Injuries can disrupt typical tissue structure and function. There are two types of wounds: acute wounds and chronic wounds. Acute wounds heal entirely in less time, while chronic wounds take longer to heal [154].
Several factors, such as wound infection, wound depth, foreign body interaction, stress, pressure, age, and congenital disease, can affect wound healing [155]. The wound healing process begins with the exudative stage, inflammation, proliferation, and finally regenerative, as shown in figure 7 [156]. The exudative stage ( figure 7(a)), also known as the hemostasis and coagulation stage, is the process of stopping blood flow within minutes by the aggregation of platelets and the formation of fibrin clots. The interaction of blood with exposed collagen and other components of the extracellular matrix causes the release of clotting factors that act as plugs by activated platelets [157].
The inflammatory stage ( figure 7(b)) releases inflammatory cells such as neutrophils and macrophages to the wound site. Foreign bodies, bacteria and damaged endogenous tissue can be cleaned by the inflammatory response mechanism [158]. Neutrophils only work in the early stages of wound healing, whereas macrophages are able to survive through all phases, from exudative to regenerative [159]. New blood vessels are formed during the proliferation phase ( figure 7(c)). In addition, the collagen fiber reinforcing synthesis process also begins in this phase. Table 3. Melting temperature (T m ), the heat of fusion (DH m ) and temperature of crystallization (T c ) of pure PVA hydrogels and PVA/CNC hydrogels determined by the DSC technique, reused with permission from Elsevier, license number: 5403010400871. Reprinted from [134], Copyright (2016), with permission from Elsevier. Furthermore, the proliferation phase also initiates the formation of epithelial cells, keratinocytes, and fibroblasts [156]. Wound healing ends with a regeneration stage ( figure 7(d)) where the dermal tissue recovers and there is an increase in the tensile strength of the scar tissue. Inflammatory cells have also been cleared from the regenerated area while collagen is renovated to increase its tensile strength. Complete wound healing may take weeks or months [160].

Number of F-T cycles
Ideally, a wound management product should be able to absorb excess wound exudate and toxins, maintain good moisture between the wound surface and the wound covering, prevent infection, prevent overheating, have good permeability, and be easily removed without aggravating the wound [161]. The scheme of using hydrogels in the wound healing process is shown in figure 8. Open wounds ( figure 8(a)) have 20 times greater water and fluid loss than normal skin and are very susceptible to bacterial invasion [118]. Applying hydrogel to an open wound (figures 8(b) and (c)) can help accelerate wound healing by exchanging moisture, thereby creating a more optimal microclimate between the wound and the wound dressing [162].
The insoluble hydrophilic structure of the hydrogel has a tremendous capacity to absorb wound exudate and allow oxygen diffusion to promote healing [163]. Hydrogels can be composited with bioactive materials so that they have antibacterial activity [79,81,164]. Hong et al successfully reported hydrogel preparations made using the F-T process for wound dressing applications [165]. They used tannic acid (TA) as an antibacterial agent and then composited it with PVA. With variations in the composition of tannic acid, PVA/TA hydrogel composite samples were frozen for 18 h at -20°C then thawed at 25°C (6 h), resulting in a hydrogel with good mechanical strength and high antibacterial activity (99.9% against S. aureus).

Drug delivery system
Hydrogels can be used in drug delivery systems because they have a three-dimensional structural network with a high-water absorption capacity. This hydrogel property is useful in the process of loading drugs and releasing drugs in a controlled manner. Drug delivery is the process of transporting pharmaceutical compounds into the body to achieve the expected therapeutic effects [166]. The high-water content present in hydrogels helps hydrophilic transport drugs. One of the physical properties of the hydrogel, which is related to its porosity, can affect the continuous drug dissolution process [154].
Two methods can carry out the process of loading the drug onto the hydrogel. First, the polymer is mixed with the drug, followed by a chemical and physical cross-linking process. The second method is the hydrogel  -Uniformity and an interconnected porous structure.
-Higher swelling ratio when CH content was increased (up to 1000%).
-Incorporation of Ag in the hydrogel network resulted in higher tensile strength and higher elongation at break (Young modulus of 14.28 MPa).
-Absorption of serum protein was increased with the increase of CH.
-Adding CeO 2 NPs increased the porosity of hydrogels (from 81 to 90%). -There was a significant increase in the swelling value after incubating it for up to 5 h, then the value tended to be constant. The added CeO2-NPs to the hydrogel structure caused an increase in swelling value, although it was not significant.
--Cells can be viable and multiply on both hydrogels (0.5% and 1%) and do not show significant toxicity. -Hydrogel with 0.5% CeO 2 NPs can effectively inhibit MRSA growth compared to hydrogels with 0% and 1%.
Human dermal fibroblasts [146] PVA/n-HA/ HACC A double network hydrogel was prepared with a variation in n-HA content (0.4, 0.6, or 0.8 g). Freeze at −20°C for 12 h and thawing at 25°C for 12 h (3 cycles).
-The DN hydrogel microstructure became denser, and the pore size decreased sharply. Meanwhile, the presence of HA did not have a vital effect on the DN hydrogel microstructure. -The sample showed a low swelling value as expected for a cartilage repair application.
-DN hydrogel with HA presence showed excellent mechanical properties (ultimate tensile strength up to 2.70 MPa and compressive strength almost 75 MPa). Without HA, the mechanical properties slightly decreased. -The presence of HA causeed the hydrogel sample to have a higher residual mass during the TG/DTG test.
-Good hydrogel ability to promote osteoblast cell growth. -Cell viability was increased with the increasing HA concentration.
The addition of COL concentrations in the hydrogel sample caused an increase in hydrophilic groups such as amino ( -NH2) and carboxyl groups ( -COOH). Thus, the hydrogel composite with 3% COL content had good Hydrogel samples with 1% COL content had good crystallinity levels so that their mechanical properties were also better (tensile strength reached 1.62 MPa and elongation at break reached 532%).   The addition of the number of F-T cycles increased the syneresis value of K-carrageenan gel. The network or structure of K-carrageenan was easily affected by the formation of ice crystals so that there was a highwater separation during the thawing process.
Food industry [148] PVA/GLE The freezing temperature at −25°C for 20 h, thawing at 24°C for 4 h for 6 cycles, and different GLE fractions.
-The morphology of the porous hydrogel was observed, which increased in size with the addition of the GLE fraction. -The ability to swell reaches 207%, which has potential as a wound dressing material.
-An increase in the modulus of elasticity reaches 1000 times the initial value at a strain ratio of 0%-95% and 15 times at a strain ratio of 50%-95%. -Shows a decrease in the degree of crystallinity of the hydrogel when the concentration of GLE is increased.
-Wound dressing [79] ALG/CS-G ALG solution was made at 2% and then added CS-G with concentration variations ranging from 0.5% to 2%. Posaconazole (POS) at 2% was added to the matrix to observe the drug release characteristics. Freeze at −20°C for 18 h and thawing at RT for 6 h, repeated for 3 cycles.
The compact structure and homogeneous pores were observed.
-Hydrogel demonstrated good mechanical properties that are suitable for skin applications with higher values of firmness, compressibility, compactness, and tackiness.
-In the thermograms, no POS temperature shift was observed, indicating no interaction between drug and polymer for all systems.
It was observed that all the prepared formulations had a similarity factor f2 < 50 and a difference factor f1 > 15, indicating that the F-T technique significantly affected the POS release profile. In addition, alginate/chitosan glutamate Drug delivery [149]  -The internal morphology observed using SEM did show some phase separation between the liquid and solid phases in the hydrogel composite. -The degree of swelling of the composite hydrogel showed an average value above 200%, which indicated a good ability to absorb wound exudate.
-The compression modulus of composite hydrogels tends to decrease with increasing PCE fraction; this is due to the presence of an ethanol fraction that does not freeze during the F-T process.
-PCE was stable in the range of room temperature, to body temperature as no significant mass decomposition was observed.
Composite hydrogel has antibacterial activity against S. aureus and P. aeruginosa. The release of PCE from the hydrogel followed a pseudo-Fickian diffusion model.
Wound dressing [81] CS/metal ions (Ag + , Ca 2+ , Zn 2+ , Cu 2+ ) 0.45% CS was loaded with different combinations of ions and their respective concentrations, then placed in the middle of medical gauzes. Freezing at −20°C overnight and thawing at RT.
The SEM results confirmed the intertwined hydrogel network that was lyophilized and filled in the mesh holes of the gauze, indicating that the CS/M n+ hydrogel was successfully coated on the gauze after the F-T process.
-In vitro studies show that gauze can temporarily release some metal ions on demand, and the released metal ions show effectiveness in killing bacteria and accelerating cell migration. In vivo studies reveal that metal ion-filled gauze can efficiently promote the healing of infected wounds Chronic and infected wound healing formed by soaking in a medicinal solution [167]. However, the drug delivery system using hydrogel media has difficulty loading hydrophobic drugs and tends to have low tensile strength [168]. The way to increase the efficiency of loading hydrophobic drugs into the hydrogel matrix is by combining molecules with the ability to form inclusion complexes and incorporate hydrophobic groups [169]. Printing of hydrogel supermolecules can increase drug loading efficiency without affecting the drug release kinetics [170].
As shown in figure 9(a), the polymer can be composited with nanoparticles to form a hydrogel superstructure for controlled drug delivery applications [171,172]. Nanoparticles have been developed very massively and have the advantage of being a type of drug delivery encapsulation that can load small molecules, peptides, nucleic acids, proteins, and so on [173,174]. The drug release characteristics of the nanoparticleshydrogel can be modified by varying the components of the nanoparticles [175].
Besides, the variations in hydrogel synthesis parameters by the F-T method can affect drug release characteristics. The greater number of F-T cycles causes the density of the cross-linking to increase so that the permeability of the matrix decreases. Thus, the release of the drug is slower [176]. Previously, Edikresnha et al had observed the release behavior of PVA hydrogels loaded by Piper crocatum extract (freeze at-20°C for 20 h, thaw at 27°C for 4 h, 6 cycles) [81]. They found that within the initial 20 min, there was a short-term burst release with nearly 13%−20% of the extract released, and even after 3 h, the extract release was still below 50%. The reason may be that the extract is not well encapsulated in the PVA matrix.
Moreover, aspects of controlled release need to be considered in the design of drug delivery systems. The main idea of controlled drug release is to treat the symptoms in the desired and targeted area [177]. Considering that the human digestive tract has a unique microenvironment, hydrogels that can respond to changes in stimulus (particularly, changes in acidity levels) are needed as a controlled drug delivery medium [178]. Currently, PVA-based hydrogels have been shown to exhibit pH-responsive properties [179,180]. However, pure PVA has the disadvantage that it is not sensitive to changes in pH [181]. Therefore, some researchers combine it with PAA to produce hydrogels with pH-responsive properties [182,183]. Nevertheless, PAA tends to cause irritation, thus limiting its use in the biomedical field.
Interestingly, Xie and co-workers in their research succeeded in developing a pH-sensitive PVA/sodium alginate hydrogel with optimized F-T process parameters (freeze at-18°C for 22 h, thaw at RT for 2 h) [184]. As a model of drug release, they used chloramphenicol, which was tested under conditions similar to the microenvironment of the human digestive tract. In figure 9(b), the cumulative release pattern of 2), confirms that the hydrogel synthesized with 4 F-T cycles and 50% sodium alginate exhibits sensitive behavior to changes in pH.
Furthermore, the effect of the sodium alginate composition in the hydrogel matrix on the release characteristics is shown in figure 9(c). The addition of sodium alginate is known to reduce the cumulative release of chloramphenicol from the hydrogel matrix, making it a potential candidate for controlled drug delivery systems. The main reason may be that the incorporation of sodium alginate into the PVA hydrogel matrix causes a decrease in the pore size, which in turn affects the drug release behavior [185]. Increasing the F-T cycle from 2 to 6 led to a reduction in the cumulative release of chloramphenicol from 83.5% to 76.6% (figure 9(d)). The reason is that the density of the hydrogel crosslinks increases with increasing F-T cycles, as previously described [176,184].

Tissue engineering
Tissue engineering is interdisciplinary research that combines several fields, such as materials science, cell biology, and engineering, to repair damaged tissue [186]. Tissue engineering aims to mimic the extracellular matrix (ECM) for tissue regeneration. The hydrogel can be used for tissue engineering, including skin regeneration, cartilage repair, bone, and vascular scaffold [7,147,187,188]. Several methods used in the manufacture of hydrogel scaffolds include lyophilization, emulsification, solvent washing, foaming gas release, photolithography, microfluidics, micro printing, and 3D printing [189].
Requirements that must be possessed by scaffolding materials include being biocompatible, biodegradable, sterile, and having a structure that supports cell attachment, cell proliferation, and differentiation [190]. Hydrogel compatibility and degradability are needed to see their interaction and characteristics with cells when applied to tissue regeneration sites [191]. Previous research by Engler's group has successfully investigated the behavior of cells under two-dimensional (2D) hydrogel conditions [192]. Several determinants of cell behavior include the level of stiffness, ligand density, and matrix porosity of the hydrogel. Huebsch et al initiated the investigation of cell behavior in 3-dimensional (3D) hydrogel conditions with variations in the level of stiffness (2.5-110 kPa) [193]. The application of the 3D hydrogel to the subchondral bone interface is illustrated in figure 10(a).
Thangprasert et al in their research, have reported on gelatin/PVA hydrogel fabrication using the F-T method for tissue engineering applications [119]. The F-T fabrication parameters, such as polymer concentration, freezing temperature, and the number of F-T cycles, greatly affect the resulting hydrogel properties. Based on their results, differences in gelatin and PVA composition led to differences in swelling properties, degradation rates, pore sizes, and biological properties (the best properties belong to the gelatin/PVA 30:70 composition). Nevertheless, the mechanical properties of the hydrogel are not sufficient to support the subchondral bone interface. However, the ALP activity test, calcium content, and alizarin red confirmed the osteogenic potential and biocompatibility of the fabricated hydrogel samples.
Fortunately, these poor mechanical properties can be improved by combining PVA with extracellular matrix components such as collagen. When bonded to PVA, collagen with a triple helical structure and a high concentration of hydrophilic groups can provide better mechanical properties [194]. In addition, collagen is able to show remarkable biocompatibility, which supports cell adhesion, proliferation, and differentiation in articular cartilage [195]. Articular cartilage plays a role in reducing friction under high loads in activities of daily movement [196].
As a natural tissue at the ends of bones in synovial joints, articular cartilage is often damaged by arthritis, sports trauma, or other conditions. When such damage occurs, the ability to repair itself is limited given the lack of blood vessels, lymphatic vessels, and nerves in the area [197]. For this reason, the incorporation of collagen and PVA in the form of hydrogels was offered as candidate articular cartilage materials [198].
Recently, a 15% PVA combined with 0.05% hyaluronic acid (HA) and collagen (COL) whose concentration varied from 0 to 3%, was synthesized into a hydrogel with the following parameters, freeze at-20°C for 8-10 h, thaw at RT for 2 h, 10 cycles [72]. As shown in figure 10(b), the F-T treatment induces polymer chains to form hydrogen bonds with each other (intramolecular, intermolecular I and intermolecular II). The F-T treatment and the increase in COL content in the hydrogel matrix resulted in a porous structure due to the increased number of hydrogen bonds. This characteristic allows for the availability of free space for water storage, which in turn can be useful for reducing friction.
Moreover, to assess the feasibility of the F-T hydrogel composite as articular cartilage, a friction lubrication test scheme was carried out ( figure 10(c)). When sliding loads are applied (loading rate 10 mm min −1 ), the soft polymer chain structure (PVA/HA/COL) can reduce the friction coefficient by up to 50% compared to pure PVA ( figure 10(d)). The decrease in the degree of cross-linking in the longer HA chains and their lubricating effect contributed to the increase in hydrogel softness and the decrease in the coefficient of friction, respectively.
It should be noted that the hydrogel is a permeable three-dimensional matrix, which allows the liquid within to be extruded when subjected to sliding loads. The liquid then acts as a lubricant for the friction pairs, and as evidenced by the increase in water retention (due to a more porous structure), the friction coefficient also decreases significantly. However, COL itself is known to have low mechanical properties, and an increase in COL content does not necessarily improve these mechanical properties [199]. From figure 10(e), it is known that the optimum tensile strength and elongation at break are shown in hydrogel composites containing 1% COL, which are 1.62 MPa and 532%, respectively [72]. Therefore, modifying the dual network structure (DN) as discussed in section 3.3, is offered to improve the mechanical properties of the hydrogel.
Gan et al in their research succeeded in developing a DN hydrogel based on poly(vinyl alcohol)-(nano hydroxyapatite)/(2-hydroxypropyltrimethyl ammonium chloride chitosan) abbreviated as PVA-HA/HACC-Cit, using the F-T method [147]. They prepared the first network of PVA-HA/HACC hydrogel under the following conditions, freeze at-20°C for 12 h, thaw at 25°C for 12 h, 3 cycles, then soak in Na 3 Cit solution to obtain a second network.
From figure 10(f), it is known that the double hydrogel network structure can increase the tensile strength and elongation at break from 0.29 MPa and 345% to 2.95 MPa and 710%. Likewise, toughness and compressive modulus show excellent mechanical properties as well. While the average coefficient of friction for all DN hydrogels (figure 10(g)) is below 0.10, which is lower than the PVA/HA/COL composite hydrogels. Moreover, excellent fatigue resistance and recovery properties have been demonstrated due to its unique dual structure. The presence of HA in the hydrogel matrix provides high tribological properties and biocompatibility, as evidenced by cell viability and proliferation assays [147]. Thus, this dual network hydrogel composite is very promising for articular cartilage applications.

Conclusion and future direction
Hydrogel is a three-dimensional material that can swell up to 1000 times its dry weight. The basic materials for hydrogel are natural polymers (for example, alginate, chitosan, gum, starch, gelatin, and cellulose) and artificial polymers (PVA, PVP, PEG, PANI, and PAA). Hydrogel synthesis can be carried out by the chemical crosslinking method (covalent bonding) and the physical cross-linking method (non-covalent).
F-T is a type of physical cross-linking method that utilizes the crystallization process of precursor solutions and hydrogen bond interactions between polymer chains to produce hydrogels. When selecting the polymer material used in the F-T method, it is necessary to consider the structure and use of the appropriate solvent. Being a hydrophilic polymer with a large amount of -OH, -COOH, and -NH 2 groups in its chain structure provides the advantage that, during the freezing process, the polymer can form chain associations through hydrogen bonding. Meanwhile, the use of solvents has a significant effect on the formation of hydrogels, considering that each of the above polymers has a different solubility. The F-T method can only be carried out if the polymer has completely dissolved so that the polymer chains with the hydrophilic functional group can crosslink each other. Thus, the use and selection of solvents according to the type of polymer must be carried out properly.
The advantage of the F-T method is that it is easy to fabricate and does not involve cross-linking chemicals, which sometimes have toxic properties. The hydrogel properties can be influenced by the parameters of the solution and the parameters of the synthesis process, such as polymer molecular weight, the solution concentration, freezing and thawing temperatures, F-T duration, and the number of F-T cycles.
The properties of the hydrogel can be adjusted according to the actual requirements of the application. Some of the most studied hydrogel properties are morphology (porosity), swelling degree, degradability, mechanical properties, thermal properties, and biocompatibility. The hydrogel produced by the F-T method has good biocompatibility. It is very supportive when applied to the biomedical field, e.g., wound dressing, drug delivery systems, and tissue engineering.
However, several challenges still need to be addressed for further development of the F-T method to yield superior hydrogel. Further study on the F-T method to produce hydrogels with the desired properties for specific applications is still required. Deep research and optimization on variations in solution and process parameters can be carried out to support this goal. For example, for wound dressing applications, hydrogels with good mechanical properties are needed. Process parameters and solution parameters are varied, which allows the formation of cross-links at the molecular level to produce an elastic, stretchable, and not easily broken hydrogel.
As we mentioned in the previous section, one of the leading users of the hydrogel product is the biomedical sector. In the biomedical field, toxicity is a crucial matter, so hydrogel-based materials that are used in this field must be free from toxic substances. Both the polymer and the solvent must satisfy this requirement. Fully natural polymers are preferred to produce hydrogels for medical applications.
Further research is still needed to explore the various sources of biopolymers and to examine their qualifications to obtain the desired hydrogels. However, hydrogel developments based on polymer blends composed of natural and synthetic ones are still worthwhile to be carried out as long as the quantity of the synthetic polymers is in the tolerance range that makes them safe to use for biomedical applications.
Many researchers have also tried to compose polymers with endemic natural materials originating from various parts of the world with antibacterial and antioxidant activity. However, the obstacle is that not all these natural materials can be composited, so further preparation is needed to investigate the properties of these natural materials before they are composited. Another obstacle in developing the F-T method for mass production on an industrial scale is that it has not guaranteed the hydrogel's properties produced on an industrial scale. Modifications and comprehensive studies are needed to observe the effect of process parameters and solution parameters on the properties of the hydrogel produced after up-scaling.
Additionally, the number of cycles in the F-T method is related to the use of energy for the freezing and thawing processes. Hence, simulation and optimization on the determination of the minimum cycle that can result in hydrogels with optimum properties for certain applications still need to be carried out.